U.S. patent number 5,217,019 [Application Number 07/815,068] was granted by the patent office on 1993-06-08 for apparatus and method for continuously monitoring cardiac output.
This patent grant is currently assigned to Abbott Laboratories. Invention is credited to Timothy J. Hughes.
United States Patent |
5,217,019 |
Hughes |
June 8, 1993 |
Apparatus and method for continuously monitoring cardiac output
Abstract
A method and apparatus for continuously monitoring cardiac
output based upon the phase shift between an in put signal and a
temperature signal indicative of the change in temperature of blood
leaving the heart. In a first preferred embodiment of a cardiac
output monitoring system (10), a catheter (14) is provided with an
electrical resistance heater (22). An electrical current having a
sinusoidal wave form with a period of from 30 to 60 seconds is
applied to the heater, causing power to be dissipated into the
blood within a patient's heart (12). A temperature sensor (24)
disposed near a distal end of the catheter produces a signal
indicative of the temperature of blood leaving the heart. The
temperature signal and the signal corresponding to the electrical
power dissipated in the heater (an input signal) are filtered at a
frequency .omega. corresponding to the frequency of the applied
electrical current, i.e., the frequency of the input signal. The
amplitude of the input power, the amplitude of the temperature
signal, and their phase difference are used in calculating cardiac
output. In another embodiment, a temperature conditioned saline
solution ( 84) is circulated through catheter (14') in a closed
loop, so that it flows through a heat exchanger (60) disposed
within the heart. The fluid is circulated through the catheter on a
periodic basis, providing a periodic input signal. The temperature
signal produced by the temperature sensor on the catheter distal
end and power dissipated to or absorbed from the blood by the heat
exchanger comprise the two signals from which the cardiac output is
determined as described above. The determination of cardiac output
is also corrected for the time constant of the catheter/heater (or
heat exchanger) and of the temperature sensor.
Inventors: |
Hughes; Timothy J. (Palo Alto,
CA) |
Assignee: |
Abbott Laboratories (Abbott
Park, IL)
|
Family
ID: |
25216763 |
Appl.
No.: |
07/815,068 |
Filed: |
December 27, 1991 |
Current U.S.
Class: |
600/481; 600/508;
600/526; 600/549 |
Current CPC
Class: |
A61B
5/028 (20130101) |
Current International
Class: |
A61B
5/028 (20060101); A61B 5/026 (20060101); A61B
005/02 (); A61B 005/00 () |
Field of
Search: |
;128/668,691,692-695,713,736 ;73/204.17,204.23,204.25 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
Other References
Philip, James H., et al., "Continuous Thermal Measurement of
Cardiac Output," IEEE Transactions on Biomedical Engineering, vol.
BME-31, No. 5, May 1984, pp. 393-400..
|
Primary Examiner: Green; Randall L.
Assistant Examiner: Burke; Elizabeth M.
Attorney, Agent or Firm: Christensen, O'Connor, Johnson
& Kindness
Claims
The embodiments of the invention in which an exclusive property or
privilege is claimed are defined as follows:
1. Apparatus for continuously monitoring a cardiac output of a
heart, comprising:
(a) a catheter having a plurality of lumens that extend generally
between a proximal end and a distal end of the catheter, the distal
end of the catheter being insertable into a heart through a
cardiovascular system;
(b) means for supplying a periodically varying, temperature
modifying input signal to a portion of the catheter that is spaced
apart from its distal end;
(c) a blood temperature sensor disposed adjacent the distal end of
the catheter, said temperature sensor being provided to produce a
blood temperature signal that is indicative of a temperature of
blood flowing from the heart;
(d) means for determining power transferred by the temperature
modifying input signal, said means for determining producing a
corresponding periodically varying power signal that is indicative
of said power transferred, which corresponds to the input
signal;
(e) a phase comparator for determining a difference in phase
between the periodically varying power signal and the periodically
varying temperature signal; and
(f) control means for determining the cardiac output of the heart
as a function of the power signal, the blood temperature signal,
and the difference in phase between said signals.
2. The apparatus of claim 1, wherein the means for supplying
comprise a source of an electrical current connected by a plurality
of leads to a resistor disposed in the portion of the catheter
spaced apart from its distal end, said input signal comprising a
periodically varying electrical current that is applied to heat the
resistor and any blood around the resistor.
3. The apparatus of claim 2, wherein the electrical current flowing
through the resistor dissipates power and wherein the means for
determining power transferred comprise means for determining the
power dissipated in the resistor by the electrical current flowing
through it.
4. The apparatus of claim 1, wherein the means for supplying
comprise a pump that delivers a temperature-conditioned fluid
through a closed loop fluid flow path defined by the lumens of the
catheter, said pump cycling on and off periodically at a predefined
frequency.
5. The apparatus of claim 4, wherein the means for determining
power transferred comprise a first temperature sensor that monitors
a temperature of the temperature-conditioned fluid pumped into the
catheter, a second temperature sensor that monitors a temperature
of the temperature-conditioned fluid as it returns from the heart,
and means for determining a rate of flow of said fluid, the control
means determining the power transferred as a function of the
difference in temperature of the temperature-conditioned fluid
monitored by the first and the second temperature sensors, and of
the rate of flow of the temperature-conditioned fluid in the
catheter.
6. The apparatus of claim 4, wherein the means for supplying
further comprise a fluid chiller that cools the temperature
conditioned fluid substantially below a temperature of blood
entering the heart.
7. The apparatus of claim 4, wherein the fluid is heated
substantially above a normal temperature of blood entering the
heart.
8. The apparatus of claim 1, wherein the cardiac output is defined
by an equation as follows:
where:
CO=the cardiac output;
P(.omega.)=the power transferred by the input signal, which varies
at an angular frequency .omega.;
.DELTA..PHI.=the difference in phase between the power signal and
the blood temperature signal;
T(.omega.)=the blood temperature indicated by the blood temperature
signal, which varies at the angular frequency .omega.; and
Cb=a specific heat times density constant for the blood.
9. The apparatus of claim 1, further comprising bandpass filter
means for filtering the power signal and the blood temperature
signal to filter out frequencies different from a frequency at
which the input signal periodically varies.
10. The apparatus of claim 1, further comprising means for
compensating for a phase shift error in the blood temperature
signal due at least in part to a thermal mass of the catheter.
11. The apparatus of claim 1, wherein the control means compensate
for an attenuation of the blood temperature signal by the catheter
and the blood temperature sensor, in determining the cardiac
output.
Description
FIELD OF THE INVENTION
This invention generally pertains to apparatus and a method for
monitoring the volumetric output of a heart, and more specifically,
to apparatus and a method for making this determination by using an
injectateless technique that changes the temperature of blood in
the heart.
BACKGROUND OF THE INVENTION
Cardiac output, the volumetric rate at which blood is pumped
through the heart, is most often determined clinically by injecting
a bolus of chilled saline or glucose solution into the right
auricle or right ventricle through a catheter. A thermistor
disposed in the pulmonary artery is used to determine a
temperature-time washout curve as the chilled injectate/blood
mixture is pumped from the heart. The area under this curve
provides an indication of cardiac output. Although this
thermo-dilution method can give an indication of cardiac output at
the time the procedure is performed, it cannot be used for
continuously monitoring cardiac output. Moreover, the frequency
with which the procedure is performed is limited by its adverse
effects on a patient, including the dilution of the patient's blood
that occurs each time the chilled fluid is injected. In addition,
the procedure poses an infection hazard to medical staff from blood
contact, and to the patient, from exposure to possibly contaminated
injectate fluid or syringes.
Alternatively, blood in the heart can be chilled or heated in an
injectateless method by a heat transfer process using a
temperature-conditioned fluid that is pumped in a closed loop,
toward the heart through one lumen within the catheter and back
through another lumen. The principal advantages of using such a
non-injectate heat transfer process to change the temperature of
blood are that the blood is not diluted, and the temperature
differential between the blood and the heat exchanger is much less
compared to the temperature differential between an injectate fluid
and blood in the typical thermal dilution procedure.
U.S. Pat. No. 4,819,655 (Webler) discloses an injectateless method
and apparatus for determining cardiac output. In Webler's preferred
embodiment, a saline solution is chilled by a refrigeration system
or ice bath and introduced into a catheter that has been inserted
through a patient's cardiovascular system into the heart. The
catheter extends through the right auricle and right ventricle and
its distal end is disposed just outside the heart in the pulmonary
artery. A pump forces the chilled saline solution through a closed
loop fluid path defined by two lumens in the catheter, so that heat
transfer occurs between the solution and blood within the heart
through the walls of the catheter. A thermistor disposed at the
distal end of the catheter monitors the temperature of blood
leaving the heart, both before the chilled fluid is circulated
through the catheter to define a baseline temperature, and after
the temperature change in the blood due to heat transfer with the
chilled saline solution has stabilized. Temperature sensors are
also provided to monitor both the temperature of the chilled saline
solution at or near the point where it enters the catheter (outside
the patient's body) and the temperature of the fluid returning from
the heart. In addition, the rate at which the chilled solution
flows through the catheter is either measured or controlled to
maintain it at a constant value. Cardiac output (CO) is then
determined from the following equation: ##EQU1## where V.sub.I
equals the rate at which the chilled fluid is circulated through
the catheter; .DELTA.T.sub.I equals the difference between the
temperature of the chilled fluid input to the catheter and the
temperature of the fluid returning from the heart; .DELTA.T.sub.B
equals the difference between the temperature of the blood leaving
the heart before the chilled fluid is circulated and the
temperature of the blood leaving the heart after the chilled fluid
is circulated (after the temperature stabilizes); and C is a
constant dependent upon the blood and fluid properties. The patent
also teaches that the fluid may instead be heated so that it
transfers heat to the blood flowing through the heart rather than
chilled to absorb heat.
U.S. Pat. No. 4,819,655 further teaches that the cardiac monitoring
system induces temperature variations in the pulmonary artery that
are related to the patient's respiratory cycle and are therefore
periodic at the respiratory rate. Accordingly, Webler suggests that
the signal indicative of T.sub.B ' (the temperature of the chilled
blood exiting the heart) should be processed through a Fourier
transform to yield a period and amplitude for the respiratory
cycle, the period or multiples of it then being used as the
interval over which to process the data to determine cardiac
output.
Another problem recognized by Webler is the delay between the times
at which circulation of the chilled fluid begins and the
temperature of the blood in the pulmonary artery reaches
equilibrium, which is caused by the volume of blood surrounding the
catheter in the right ventricle and in other portions of the heart.
The patent suggests introducing a generally corresponding delay
between the time that temperature measurements are made of the
blood before the chilled fluid is circulated and after, for
example, by waiting for the .DELTA.T.sub.B value to exceed a level
above that induced by respiratory variations. However, for a
relatively large volume heart and/or very low cardiac output, the
T.sub.B ' data do not reach equilibrium in any reasonable period of
time. The quantity of blood flowing through the large volume heart
represents too much mixing volume to accommodate the technique
taught by Webler for processing the data to determine cardiac
output. As a result, the measurement period for equilibrium must be
excessively long to reach equilibrium, thereby introducing a
potential error in the result due to either a shift in the baseline
temperature of the blood or changes in the cardiac output. For this
reason, the technique taught by Webler to determine cardiac output
using the data developed by his system is not practical in the case
of large blood volumes in the heart and/or low cardiac outputs.
The technique disclosed by Webler also assumes that all of the
energy absorbed by a chilled fluid (or lost by a heated fluid)
represents heat transferred between the fluid and the blood in the
heart. This assumption ignores the heat transfer that occurs
between the fluid and the mass of the catheter. A somewhat smaller
source of error arises due to the energy required to change the
temperature of the small thermal mass of the thermistor bead that
monitors the temperature of blood leaving the heart. For long
measurement periods, these errors can generally be ignored. In
addition, if the thermistor bead is selected to have a very small
mass and fast response time, its error contribution may be
insignificant. However, as the measurement period becomes shorter,
the effect of these error sources becomes increasingly more
important.
Instead of cooling (or heating) the blood in the heart by heat
transfer with a circulating fluid to determine cardiac output, the
blood can be heated with an electrical resistance heater that is
disposed on a catheter inserted into heart. The apparatus required
for this type of injectateless cardiac output measurement is
significantly less complex than that required for circulating a
fluid through the catheter. An electrical current is applied to the
resistor through leads in the catheter and adjusted to develop
sufficient power dissipation to produce a desired temperature rise
signal in the blood. However, care must be taken to avoid using a
high power that might damage the blood by overheating it. An
adequate signal-to-noise ratio is instead preferably obtained by
applying the electrical current to the heater at a frequency
corresponding to that of the minimum noise generated in the
circulatory system, i.e., in the range of 0.02 through 0.15 Hz.
U.S. Pat. No. 4,236,527 (Newbower et al.) describes such a system,
and more importantly, describes a technique for processing the
signals developed by the system to compensate for the above-noted
effect of the mixing volume in the heart and cardiovascular system
of a patient, even one with a relatively large heart. (Also see J.
H. Philip, M. C. Long, M. D. Quinn, and R. S. Newbower, "Continuous
Thermal Measurement of Cardiac Output," IEEE Transactions on
Biomedical Engineering, Vol. BMI 31, No. 5, May 1984.)
Newbower et al. teaches modulating the thermal energy added to the
blood at two frequencies, e.g., a fundamental frequency and its
harmonic, or with a square wave signal. Preferably, the fundamental
frequency equals that of the minimal noise in the cardiac system.
The temperature of the blood exiting the heart is monitored,
producing an output signal that is filtered at the fundamental
frequency to yield conventional cardiac output information. The
other modulation frequency is similarly monitored and filtered at
the harmonic frequency, and is used to determine a second variable
affecting the transfer function between the injection of energy
into the blood and the temperature of the blood in the pulmonary
artery. The amplitude data developed from the dual frequency
measurements allows the absolute heart output to be determined,
thereby accounting for the variability of fluid capacity or mixing
volume.
Newbower et al. does not address correcting for errors due to the
thermal mass of the catheter and the thermistor bead used to
monitor the temperature of blood leaving the heart. Furthermore,
the technique taught in Newbower et al. requires matching the dual
frequency data to a predefined curve using a best fit algorithm, to
determine the absolute cardiac output. Accordingly, the results are
not as accurate as may be desired, particularly in the presence of
noise.
It is preferable that a non-injectate method for determining
cardiac output be based on measured output data processed using a
technique that does not require fitting the output data to a curve.
Cardiac output should also be determined by a method that
compensates for the mixing volume of the heart, regardless of its
relative size, and also compensates for the thermal mass of the
catheter and the thermistor bead used to produce the output signal.
The foregoing aspects and many of the attendant advantages of the
present invention will become more readily appreciated as the same
becomes better understood by reference to the following detailed
description, when taken in conjunction with the accompanying
drawings.
SUMMARY OF THE INVENTION
In accordance with the present invention, apparatus are provided
for continuously monitoring a cardiac output of a heart. The
apparatus include a catheter having a plurality of lumens that
extend generally between a proximal end and a distal end. The
distal end of the catheter is insertable into the heart through a
cardiovascular system. Means are also included for supplying a
periodically varying, temperature modifying input signal to a
portion of the catheter that is spaced apart from its distal end. A
blood temperature sensor is disposed adjacent the distal end of the
catheter and produces a blood temperature signal that is indicative
of a temperature of blood flowing from the heart. Means are
operative to determine power transferred by the temperature
modifying input signal, producing a corresponding periodically
varying power signal that is indicative of the power transferred. A
phase comparator determines a difference in phase between the
periodically varying power signal and the periodically varying
temperature signal. Control means then determine the cardiac output
of the heart as a function of the power signal, the blood
temperature signal, and the difference in phase between said
signals.
Preferably, the means for supplying the periodically varying
temperature modifying signal comprise a source of an electrical
current connected by a plurality of leads to a resistor disposed on
a portion of the catheter that is spaced apart from its distal end.
The input signal comprises a periodically varying electrical
current that is applied to heat the resistor and any blood around
the resistor. The means for determining power transferred comprise
means for determining the power dissipated in the resistor by the
electrical current flowing through it.
Alternatively, the means for supplying the periodically varying
temperature modifying signal comprise a pump that delivers a
temperature-conditioned fluid through a closed loop fluid flow path
defined by the lumens in the catheter. The pump cycles on and off
periodically at a predefined frequency. For this embodiment, the
means for determining power transferred comprise a first
temperature sensor that monitors the temperature of the
temperature-conditioned fluid pumped into the catheter, a second
temperature sensor that monitors the temperature of the
temperature-conditioned fluid as it returns from the heart, and
means for determining the rate of flow of said
temperature-conditioned fluid. The control means determine the
power transferred as a function of the difference in temperature of
the temperature-conditioned fluid monitored by the first and the
second temperature sensors, and the rate of flow of the
temperature-conditioned fluid in the catheter.
Instead of being chilled, the temperature-conditioned fluid may be
heated substantially above a normal temperature of blood entering
the heart.
The cardiac output is defined by the following equation:
where:
CO=the cardiac output;
P(.omega.)=the power transferred by the input signal, which varies
at a frequency .omega.;
.DELTA..PHI.=the difference in phase between the power signal and
the blood temperature signal;
T(.omega.)=the blood temperature indicated by the blood temperature
signal, which varies at the frequency .omega.; and
Cb=a specific heat times density constant for the blood.
The apparatus further comprises filter means for filtering the
power signal and the blood temperature signal to remove frequencies
different from the frequency at which the input signal periodically
varies. In addition, the control means compensate for an
attenuation of the blood temperature signal by the catheter and the
blood temperature sensor, in determining the cardiac output.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is a block diagram of a first embodiment of the present
invention, illustrating the disposition of a catheter and
electrical resistance heater within a human heart that is cut away
to more clearly show the right auricle, ventricle, and pulmonary
artery;
FIG. 2 is a cut away view of a human heart, showing the disposition
of a catheter through which a temperature-conditioned fluid is
circulated to change the temperature of the blood within the
heart;
FIG. 3 is a block diagram of a cardiac output measurement system
used in connection with a noninjectate system that changes the
temperature of blood in the heart by heat exchange with a fluid
circulated through a catheter in a closed loop; and
FIG. 4 is a flow chart showing the logical steps used in
determining cardiac output in accordance with the present
invention.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT
A first embodiment of a cardiac output monitoring system in
accordance with the present invention is shown generally in FIG. 1
at reference numeral 10. A human heart 12 is schematically
illustrated in this figure, with a portion of the heart cut away to
show the disposition of a catheter 14 that is inserted through a
patient's cardiovascular system and into heart 12. Catheter 14 has
a proximal end 16 and a distal end 18. A plurality of leads 20
extend longitudinally through catheter 14 (within lumens that are
not separately shown) and include leads 20a and 20b that carry an
electrical current to an electrical resistance heater 22. In the
preferred form of the invention, heater 22 comprises a coil of
insulated copper, stainless steel, nickel, or nichrome wire
approximately 12 cm in length that is wound around catheter 14
approximately 10 to 15 cm from distal end 18. Heater 22 has a
nominal resistance of from 15 to 30 ohms. Leads 20c are connected
to a temperature sensor 24, which is spaced apart from distal end
18 and generally mounted on the external surface of the catheter so
that it can readily sense the temperature of blood flowing past the
distal end as the blood is pumped from heart 12.
As shown clearly in FIG. 1, catheter 14 extends through a right
auricle 26, a right ventricle 28, and into a pulmonary artery 30 of
the patient whose cardiac output is being monitored. Adjacent
distal end 18 is disposed a balloon 32, which is inflated to float
distal end 18 upwardly from right ventricle 28 into pulmonary
artery 30. Heater 22 can be positioned entirely within right
auricle 26, or as shown, may extend from right auricle 26 into
right ventricle 28.
A regulated current supply 34 supplies a periodic electrical
current used to generate heat in a sinusoidal wave form at heater
22, at a voltage ranging from 10 to 25 volts peak amplitude.
Alternatively, a square wave current supply can be used. As the
current flows through the wire coil comprising heater 22, it
produces heat in proportion to the I.sup.2 R losses in the heater
(where I is the current and R is the resistance of the heater). The
heat produced is transferred to the blood within right auricle 26
and right ventricle 28. A current sensor 36 produces a signal
indicative of the magnitude of the electrical current flowing
through lead 20a to heater 22, and this signal is input through
leads 38 to analog-to-digital (A-D) converters 40. A second input
to A-D converters 40 is a voltage signal that indicates the voltage
developed across heater 22; this voltage signal is conveyed by a
lead 42. The third input to the A-D converters comprises the signal
indicative of the temperature of the blood leaving heart 12,
produced by temperature sensor 24, connected to leads 25, which
comprise the distal end of leads 20c. Digitized signals from A-D
converters 40 are conveyed through leads 44 to input ports (not
separately shown) on a portable computer 46.
Associated with portable computer 46 is a video display 48 on which
data defining the cardiac output of heart 12 are displayed, along
with other data and information. A keyboard 50 is connected to
portable computer 46 to provide for input and user control of the
cardiac output measurement. In addition, portable computer 46
includes a hard drive or floppy drive 52 that is used for magnetic
storage of data, test results, and programs such as the software
controlling the measurement of cardiac output. Portable computer 46
controls regulated current supply 34 by supplying control signals
transmitted through leads 54 that extend between the regulated
current supply and the portable computer.
The electrical current that energizes heater 22 to heat the blood
flowing through heart 12 is supplied either in the form of a sine
wave having a 30- to 60-second period, or as a square wave with an
energized period ranging between 15 and 30 seconds (followed by a
like duration during which no current is supplied). The power
developed by heater 22 thus represents a periodic input signal,
whereas the signal developed by temperature sensor 24 comprises an
output signal indicative of the temperature of the blood leaving
the heart. To determine power dissipated within heater 22, the
digitized signals indicative of the current flowing through the
heater and voltage drop across it are multiplied together by
portable computer 46. The power dissipated within heater 22 to heat
the blood flowing through heart 12, i.e., the peak to peak
amplitude, is therefore easily determined and is defined as the
"input signal" for purposes of the following discussion.
Accordingly, the power applied, which represents the input signal,
and the temperature of the blood exiting the heart through the
pulmonary artery, which represents the output signal, are used in
the first preferred embodiment to determine the cardiac output of
heart 12, as explained below.
An alternative embodiment for developing an input signal and an
output signal that can be used to determine the cardiac output of
heart 12 is shown in FIG. 2. In this embodiment, a catheter 14' is
used to convey a cooling or heating fluid to a heat exchanger 60
formed on the catheter, set back from its distal end so that the
heat exchanger is within right auricle 26. Two lumens (not
separately shown) within catheter 14' define a supply fluid path 62
through which a liquid cooled to a temperature well below that of
the body temperature of the patient is conveyed to heat exchanger
60, and a return fluid path 64 through which the fluid is then
returned back to a source of the fluid, external to the patient's
body. In most other aspects of its configuration and use, catheter
14' is similar to catheter 14, shown in FIG. 1. Like catheter 14,
catheter 14' includes temperature sensor 24 disposed adjacent its
distal end 18 so that it is positioned within pulmonary artery
30.
Instead of cooling a fluid to a temperature lower than the
temperature of blood entering heart 12 through catheter 14', the
fluid may be heated above the temperature of the blood so that it
transfers heat to the blood, just as heater 22 does. In either
case, whether the input signal cools the blood or heats it, the
cardiac output measurement system changes the temperature of blood
in the heart on a periodic basis so that the output signal produced
by temperature sensor 24 changes periodically in response thereto.
Furthermore, the change in the temperature of blood flowing from
the heart, i.e., the output signal, is phase shifted relative to
the input signal due to the time required to change the temperature
of the mass of blood within the right auricle and right
ventricle.
In FIG. 3, the remainder of a cardiac output measurement system 80,
which is used for circulation of a temperature conditioned fluid
(with respect to the temperature of blood entering heart 12)
through catheter 14' is illustrated schematically. Cardiac output
measurement system 80 includes a reservoir 82 (hanger bag) of a
saline solution 84. Saline solution 84 flows under the influence of
gravity through a line 86 to a pump 88. When energized for periods
of 15 to 30 seconds at a time, pump 88 forces saline solution 84
through a supply line 90, which is connected to supply fluid path
62 within catheter 14'. After the liquid flows through catheter 14'
and exchanges heat with blood within heart 12 at heat exchanger 60,
it flows back along return fluid path 64 into a return line 92.
Return line 92 passes through an external heat exchanger 96, which
reduces the temperature of the returning saline solution to ambient
temperature, e.g., 24.degree. C. Thereafter, the returning saline
solution flows back into reservoir 82 for recirculation by pump
88.
The operation of pump 88 is controlled by a pump control 98, which
is connected to the pump by leads 100 that convey signals
determining the rate at which pump 88 operates. In addition, leads
100 carry an ENABLE signal that energizes pump 88 and signals
indicative of any alarm condition, e.g., air in the line or
restriction of lines 86 or 90. Pump control 98 also receives a
signal from pump 88 indicating that the pump is running, to confirm
that fluid is being delivered to the catheter as expected.
It will be appreciated that instead of using liquid at ambient
temperature to cool the blood flowing through the heart, saline
solution 84 can be chilled to a much cooler temperature (using a
chiller coil disposed downstream of pump 88, in heat transfer
relationship with supply line 90). For example, saline solution 84
can be chilled to a lower than ambient temperature by heat transfer
with ice water at 0.degree. C.; or, a more elaborate evaporative
refrigerant chiller coil can be employed that uses a refrigerant
fluid to cool saline solution 84 as the refrigerant fluid undergoes
expansion. Similarly, it is also possible to provide heat transfer
between saline solution 84 that is circulated through catheter 14'
and a heated liquid or to provide heat from some other source so
that the saline solution entering catheter 14' is elevated in
temperature above the temperature of blood entering heart 12.
Pump control 98 is controlled by portable computer 46 so that pump
88 is enabled on a periodic basis to circulate temperature
conditioned saline solution 84 through catheter 14'. In this
embodiment, the input signal to the blood within the heart is
represented by the flow of temperature conditioning liquid through
catheter 14'. A signal applied to pump control 98 over lines 103,
which connect the pump control to the portable computer, is used to
enable the operation of pump 88. The flow of
temperature-conditioned saline solution 84 through catheter 14' is
enabled for 15 to 30 seconds, then turned off for an equivalent
interval, and this cyclic operation is continued during the
measurement of cardiac output.
A plurality of lines 102 carry signals indicative of various
temperatures to A-D converters 40, which supplies the corresponding
digitized signals to portable computer 46. Specifically, a line
102a is connected to lead 20c, and thus conveys the signal
indicative of the temperature of blood leaving heart 12 to A-D
converters 40. A lead 102b is connected to a temperature sensor 104
that produces a signal indicative of the temperature of saline
solution 84 flowing into supply fluid path 62 within catheter 14'.
Similarly, a temperature sensor 106 is connected to a lead 102c,
which conveys a signal indicative of the temperature of saline
solution 84 returning from catheter 14' into return line 92. A
plurality of fluid lines 94 are connected to other lumens within
catheter 14' and can be used to inject medication into the heart
and inflate balloon 32 during the insertion of catheter 14' into
heart 12.
As noted in the Background of the Invention, the present invention
enables cardiac output to be determined continuously rather than
intermittently (an unfortunate limitation of the conventional
injectate thermal dilution technique) and is much less prone to
noise than previous continuous cardiac output monitoring methods.
In the present invention, cardiac output is determined by portable
computer 46 following the logical steps shown in a flow chart 120,
in FIG. 4. Starting at a block 122, the temperature of blood
flowing through heart 12 is modified by applying the input signal,
e.g., by electrical current to heater 22, or by initiating the flow
of a temperature-conditioned fluid through catheter 14' so that
heat is transferred at heat exchanger 60--in either case, thereby
modifying the temperature of blood within the heart. The transfer
of heat to or from blood within heart 12 occurs at a frequency
.omega., as shown in block 122. This frequency is selected to
minimize the noise caused by patient respiration.
A dashed-line block 124 indicates that the blood heated or cooled
by the input signal mixes with the other blood in right ventricle
28 and enters pulmonary artery 30. A block 126 refers to
temperature sensor 24, which produces the signal that is indicative
of the temperature of blood exiting heart 12. With reference to a
block 128, the blood temperature T within pulmonary artery 30
comprises the output signal that is digitized by A-D converter 40.
The digitized signal indicative of the temperature of blood within
the pulmonary artery is filtered at the input frequency .omega., as
indicated in a block 130 in FIG. 4.
In the preferred embodiment, the output signal is filtered by
portable computer 46. Specifically, a discrete Fourier transform is
performed on the digitized output signal to transform the signal
from the time domain into the frequency domain. The portion of the
transformed signal at the input frequency .omega. is thus
determined and comprises a filtered output signal. By filtering the
output signal (and the input signal, as described below), noise at
other frequencies is substantially eliminated. Alternatively, an
analog bandpass filter circuit could be used to process the input
signal before it is digitized, in lieu of the discrete Fourier
transform. Other types of digital filtering could also be used to
eliminate noise components at other frequencies.
After the output signal is filtered, the amplitude of the filtered
output signal is determined, as noted in a block 132. Portable
computer 46 uses the peak to peak value of the filtered output
signal for this amplitude, represented by
.vertline.T(.omega.).vertline.. The value
.vertline.T(.omega.).vertline. is then used in a block 134 for
calculating cardiac output. Since the filtered output signal is a
periodically varying signal, it has a phase relationship that is
represented by the value .PHI..sub.out (used as described
below).
The left side of flow chart 120 is directed to the steps used in
processing the input signal. As shown in a block 138, the heating
or cooling power P, which represents the heat transferred to or
absorbed from the blood in the heart, is determined. As described
above, the heating power of heater 22 is determined from the
product of the electrical current flowing through it and the
voltage drop across the heater, as well known to those of ordinary
skill in the art.
If catheter 14' is used and heat is transferred between the
circulated saline solution and blood flowing through heart 12, the
input signal is determined as a function of: (a) the temperature
differential between the saline solution supplied to catheter 14'
and that returning from the catheter, measured at temperature
sensors 104 and 106, and (b) the saline solution flow rate provided
by pump 88. In the preferred form of the invention shown in FIG. 3,
pump 88 is set to provide a flow rate of approximately 1.5 liters
per hour when energized. The input signal (representing input power
P) is determined by portable computer 46 from the digitized signals
indicative of the saline solution temperatures at temperature
sensors 104 and 106, the flow rate of the saline solution through
the catheter (predefined or measured), and the specific heat of the
saline solution, as will be apparent to those of ordinary skill in
the art.
Portable computer 46 then filters the input signal at the input
frequency .omega., as indicated in a block 140. To filter the input
signal, the portable computer processes it with a discrete Fourier
transform, converting it from the time domain to the frequency
domain. The portion of the transformed signal at the frequency
.omega. comprises the filtered input signal. The filtered input
signal has both a phase and amplitude. In a block 142, the
amplitude of the input signal is determined and is input to block
134 as .vertline.P(.omega.).vertline.. The phase of this filtered
input signal, .PHI..sub.in, is compared to the phase of the output
signal in a block 136, producing a differential phase .DELTA..PHI.,
which is equal to the difference between .PHI..sub.in and
.PHI..sub.out. Portable computer 46 determines the differential
phase and as shown in block 134, calculates cardiac output "CO" as
follows:
In the above equation, the value Cb is the product of specific heat
and density of blood.
The volume of blood within right ventricle of heart 12, i.e., the
mixing volume, is estimated from the following expression: ##EQU2##
where .tau. is the period of the input signal. To reduce the
effects of phase noise on the determination of cardiac output, an
estimation of mixing volume can be made from Equation 4 and used in
the following relationship: ##EQU3##
The estimate of mixing volume is preferably averaged over a long
term (assuming that volume is relatively constant over the time
during which cardiac output is determined), yielding an average
mixing volume, V, which is used in Equation 5 to determine cardiac
output. The resulting determination of cardiac output from Equation
5 is therefore less sensitive to phase noise, including heart rate
variations.
When a heat signal is injected into the blood within heart 12,
either by cooling the blood or by applying heat to it, a transport
delay time is incurred before the input heat signal reaches
temperature sensor 24 in the pulmonary artery. The transport delay
time adds a phase shift that is flow rate and vessel size
dependent. The phase error due to transport delay time is defined
as: ##EQU4## where L is equal to the length of the path from the
point of which the heat signal is injected into the blood within
the heart to the point at which the temperature sensor is disposed
(in cm), R is the vessel radius (in cm), and CO is the cardiac
output in liters/second. For example, a typical phase shift would
be approximately 28.8.degree. for a path 10 cm in length, with a
rate of flow of one liter per minute, a radius of 1.6 cm, and a
period for the injection of the heat signal equal to 60
seconds.
The phase shift introduced by transport delay becomes significant
at relatively low flow rates, making accurate correction for the
mixing volume difficult. One way to address this problem is to
apply the input signal at two (or more) different frequencies,
enabling a separate estimate of transport delay phase shift and
mixing volume phase shift to be determined from the difference in
phase shift at the different frequencies.
There are two additional sources of error for which corrections can
be applied in determining cardiac output. The sources of error
relate to the time constant for the catheter and thermistor caused
by their respective thermal masses. The thermal mass of the
catheter attenuates and phase shifts the input signal, whereas the
thermal mass of temperature sensor 24 attenuates and phase shifts
the received temperature signal corresponding to the change in
temperature in the blood flowing past temperature sensor 24. The
correction used in the preferred embodiment assumes a simple
first-order system. For example, heater 22 is assumed to have a
time constant T.sub.htr (actually the time constant is for the
catheter and heater), and temperature sensor 24 to have a time
constant T.sub.sens, both of which are empirically determined.
Cardiac output is then determined from: ##EQU5## where:
.PHI..sub.htr =-ARCTAN(.omega.*T.sub.htr);
.PHI..sub.sens =-ARCTAN(.omega.*T.sub.sens);
HTR.sub.atten =COS(.PHI..sub.htr); and
SENSOR.sub.atten =COS(.PHI..sub.sens).
Equation 7 recognizes that a time delay occurs between the arrival
at temperature sensor 24 of blood having a different temperature
due to the input of a heat signal and the change in the output
signal of the temperature sensor. Similarly, the thermal mass of
the catheter/heater introduces a time delay between the application
of the input signal and the transfer of energy into the blood
around heater 22 (or heat exchanger 60). Typical time constants for
both heater 22 and temperature sensor 24 are approximately two
seconds each. Based on the assumption that the time constants for
these two elements do not vary with flow rate, amplitude errors and
thus cardiac output errors introduced from this source of error,
should be constant, dependent only on the frequency of the input
signal. Accordingly, the phase shift introduced by these time
constants should also be constant. Since the sensitivity to phase
errors increases at low flow rates and large mixing volumes, it is
important to correct for the phase shift due to the time constants
of the catheter/heater (or heat exchanger) and temperature sensor,
at large overall phase angles.
While the preferred embodiment of the invention has been
illustrated and described, it will be appreciated that various
changes can be made therein without departing from the spirit and
scope of the invention. Accordingly, it is not intended that the
scope of the present invention be in any way limited by the
disclosure of the preferred embodiment, but instead that it be
determined entirely by reference to the claims that follow.
* * * * *